1. Field of the Invention
The present invention relates to a diagnostic ultrasound apparatus and a method of diagnosis with ultrasound images capable of visualizing images of blood flow as an element in motion in an object to be imaged at a higher degree of resolution and high sensitivity, and in particular, to the apparatus and method preferable to an contrast echo technique performed with an ultrasound contrast agent injected into an object for acquiring blood flow images.
2. Description of the Related Art
A diagnostic ultrasound apparatus has a wide variety of advantages, such as, relatively compact in size, lower price, no X-ray exposure, and availability of blood flow imaging on an ultrasound Doppler technique, thereby providing an indispensable today""s imaging modality in the field of clinics.
Particularly, the blood flow imaging on the ultrasound Doppler technique, which functions strongly in finding lesions in the cardiac system and others, is called color flow mapping (CFM) or color Doppler tomography and has been standardized in almost all the diagnostic ultrasound apparatus. This color flow mapping two-dimensionally displays blood flow information in almost real time, where a flow toward an ultrasound probe is displayed in red, while a flow away from the probe is displayed in blue.
To perform such display, it is required that the same location in an object be ultrasound-scanned a plurality of times, N, to acquire a time-sequential echo signal and a velocity of blood cells at a desired depth location is detected from the echo signal on the Doppler technique. That is, a Doppler signal is derived from an amount of phase shift per unit time of a reflected signal (blood flow signal) originated from blood cells through scanning of the same location at intervals, and converted into a velocity of blood flow.
In an echo signal associated with ultrasound scanning at each time, there are mixed of a reflected wave from moving substances such as blood cells and another reflected wave from stationary substances, such as the wall of a blood vessel and tissue, which scarcely move. It is characteristic that the latter is dominant in reflected intensity, but there occurs almost no Doppler shift in the reflected wave (clutter signal) from the stationary reflection. members belonging to the latter, while the former has the Doppler shift therein. Thus a Doppler signal is extracted from the echo signal by a quadrature phase detector (comprising a mixer and an LPF), and a blood flow Doppler signal is efficiently extracted by an MTI filter which removes a clutter signal component from the Doppler signal on differences in Doppler shift amounts. The blood flow Doppler signal then experiences frequency analysis carried out using N-piece Doppler data at each depth location, producing an average of the spectrum. (Doppler frequency), dispersion, or reflection intensity from blood cells. The Doppler frequency fd is converted into a Doppler velocity vd according to this formula:
xe2x80x83vd=fdxc2x7c/(2fMxc2x7cosxcex8)xe2x80x83xe2x80x83(1),
wherein c denotes a sound velocity, fM does the frequency of a reference signal in the mixer, xcex8 does an angle made between an ultrasound beam and blood flow. This information about blood flow thus-obtained is, in general, two-dimensionally displayed on a monitor, with a B-mode image employed as a background.
A CFM mode used in performing this color flow mapping (CFM) will now be compared with a B mode in terms of resolution, S/N, dynamic range for display, aliasing frequency, realtime performance, and others.
The number of burst waves associated with transmitted ultrasound waves differs between the B mode and CFM mode. The burst wave number is defined as the number, per cycle, of ultrasound pulses having the length of a transmission repetition T0 that is the inverse of an ultrasound transmission frequency f0.
The B mode is directed to observing a tomographic image, that is, an image composed of ultrasound signals echoed from tissue. The reflected signal from the tissue is able to have a satisfactorily high S/N, because the reflected signal can be detected at fully large signal intensities within a range of ultrasound pressures which are determined with consideration of safety for an object to be diagnosed. Hence the burst wave number can be set to a lower value, such as one to two waves, and a range resolution can be increased satisfactorily, fulfilling both the S/N and range resolution.
By contrast, the CFM mode is used for observing blood flow, which corresponds therefore to a reflected signal from blood cells (blood flow signal). This blood flow signal is considerably less in signal intensity, approximately xe2x88x9240 to xe2x88x9280 dB, than that acquired from the tissue. Under the same transmission pulse condition as that in the B mode, the CFM mode provides an inferior S/N, with blood flow information substantially unavailable.
The S/N can therefore be improved by increasing the power of an ultrasound pulse to be transmitted. However, in general, since the transmitted ultrasound pressure has a limitation that has been determined with consideration of safety for an object in the B mode, it is difficult to raise the pressure any more. As a result, the number of burst waves is determined at a larger value, such as three waves or more, to enhance power of the ultrasound pulse to be transmitted. An excessively large burst wave number cause, however, range resolution to be deteriorated, thus an upper limit of the burst wave number being determined dependently on an allowed value of the range resolution.
Although the S/N of the blood flow signal can be improved in this way, the power of the blood flow signal still remain smaller by approx. a few dozes of dB than the power of a single reflected from tissue, even when the burst wave number would be raised up to the upper limit within the tolerance thereof. This causes differences in dynamic ranges for display. The dynamic range for B-mode display is large; for example, 100 dB in maximum, although the dynamic rage for the power mode displaying power under the CFM mode is small; for example, 40 dB in maximum.
The ultrasound pulse is repetitively transmitted at a pulse repetition time Tr. Thus in the velocity mode displaying Doppler velocities under the CFM mode, an aliasing phenomenon will occur, due to the sampling theory, at xc2x1fr/2 which is half the pulse repetition frequency fr=1/Tr inverted from the pulse repetition time. The values of xc2x1fr/2 are referred to as aliasing frequencies. The sign xc2x1 means that the direction is separated. From the foregoing equation (1), an aliasing velocity vr/2 corresponding to the aliasing frequencies is obtained as follows, by setting xcex8=0:
vr/2=(fr/2)xc2x7(c/2fM)xe2x80x83xe2x80x83(2).
Because c and fM are constant, the aliasing velocity vr/2 becomes constant as well. This aliasing velocity is normally displayed, for diagnosis, on a TV monitor together with a two-dimensional image indicative of blood flow information.
In the B mode, a tomographic image is obtained by performing one time transmission and reception of an ultrasound pulse along the same raster (beam) direction, while in the CFM mode, imaging is based on Doppler signals obtained by performing the transmission and reception of an ultrasound pulse a plurality of times along the same raster direction. Thus, the CFM mode is largely lowered in frame rate than the B mode. For instance, where the transmission and reception is desired to be repeated sixteen times in the same direction, the transmission and reception is required to be repeated seventeen times in total, including scanning of the B mode. If the number of frames for the B mode is ten frames per sec., the number for the CFM mode is six frames per sec., thereby reducing realtime performance.
As a countermeasure to improve the realtime performance in the CFM mode, a technique called xe2x80x9cparallel simultaneous receptionxe2x80x9d has now been in practical use, where the transmission is performed in one direction and the reception is performed simultaneously in a plurality of directions. However, this parallel simultaneous reception technique requires a transmitted beam to be spread, which reduces power transmitted to each sample position (depth position) in an object, ultimately lowering detection sensitivity. Since the safety standard regulates the transmitted power not to exceed a given value, performing the parallel simultaneous reception generally reduces the detection sensitivity. Therefore, the CFM mode, which is in charge of imaging signals of lower sensitivities, i.e., blood flow signals, has a limitation in using the parallel simultaneous reception technique. Namely, a portion to be diagnosed is confined to portions having relatively higher sensitivities, such as ventricles. The parallel simultaneous reception technique is still effective in improving realtime performance, but this technique is not a way that can always be used regardless of portions to be diagnosed.
On one hand, to diagnose tumor or ischemic heart illness, there has been a strong need that one would like to detect blood vessels as thinner as possible, like seen in detecting blood flow in tumor or blood flow of the coronary. In the conventional CFM mode, the following countermeasures to improve sensitivity was planned, in addition to enhance the fundamental performance of an apparatus by adopting higher-performance components.
Namely, the detection on the power mode has been improved recently, and is superior in sensitivity than the velocity mode. The reason is as follows.
In the velocity mode, because a flow of which velocity is almost zero or a flow intersecting an ultrasound beam becomes back when it is visualized, which means the velocity is not practically shown. In the case of the improved power mode, those types of blood flow are visualized based on the strength of their power, providing greater detectability in lower-velocity flows.
Moreover, in the velocity mode, a signal whose intensity is less than a given value is always regarded as noise and is avoided from being displayed, even if it is a blood flow signal. On one hand, according to the improved power mode, even when signal power is weak, it is displayed at a lowered luminance. And a spatial connection of signals is considered into whether the signals are from blood flow, with the result that blood flow of which sensitivity is low is easy to detect.
Though the improved power mode has owned such enhanced sensitivity, the imaging conditions of the conventional CFM mode are applied to the improved power mode. Potential ability of the improved power mode has not been shown fully yet. For example, the number of transmission burst waves still remains as conventional, thus lowering spatial resolution. As a result, a thin blood vessel is displayed as being thick one or adjoining blood vessels are displayed without being separated, only having shown poor diagnosis performance.
Under the circumstances, an evaluation of blood flow utilizing an ultrasound contrast agent has been tried recently. Since the ultrasound contrast agent (hereinafter, referred to as a contrast agent) enhances the scattering intensity of ultrasound signals when it is administered into a blood vessel of an object, it has been expected to obtain blood flow images of superior diagnosis performance with the use of the enhancement effect. For a recent few years, in particular, the contrast agent has been remarkably increased in its performance and has upgraded its contrast effect, in addition to lowered invasiveness due to the fact that the contrast agent is possible to administer from the vein, it seems that the contrast agent will become more and more popular from now. Associated with this fact, there is a need that a diagnostic ultrasound apparatus should has a function to perform diagnosis making use of the features of the contrast agent improved year by year.
In the case that blood flows of an object into which this contrast agent is injected are observed with the conventional diagnostic ultrasound apparatus, the following various problems has remained unsolved at present, to one""s regret.
The contrast agent is injected into an object to enhance the sensitivity of blood flow signals. Practically, such agent is administered from the body surface to the vein or from a catheter to the artery, and flows into each organ through the heart and/or large arteries. The primary constituent of the contrast agent consists of microbubbles having a diameter of approx. a few microns, and is greatly higher in scattering intensity than the blood cells ( for example, it is higher by a few dozes of decibels). Injecting this constant medium allows blood flow signals (in detail, echo signals that are reflected from the contrast agent flowing through blood vessels) to be enhanced largely up to a comparable level with echo signals emanated from tissue. It seems that this enhancement makes it possible to detect blood vessels whose diameters are thin or which exist deeply, which have been undetectable so far.
However, in a practical use, there has been reported that a phenomenon called xe2x80x9cbloomingxe2x80x9d occurs in which a blood flow is displayed with its picture creeping out largely, compared to the diameter of a blood vessel displayed in the B mode. FIG. 38(a) illustrates an example of an ordinary Doppler velocity image which depicts a blood vessel B, while FIG. 38(b) does an example of a Doppler velocity image in which there occurs the blooming due to the injection of a contrast agent. The blooming, when it occurs, extremely deteriorates the spatial resolution, thus practical diagnosis being almost impossible.
It is considered that the blooming will occur by the following reasons. Assume that the pulse length of an ultrasound pulse to be transmitted is assigned to a value determined by the number of burst waves, M (positive integer) and a transmission frequency f0. The pulse length at a time when being outputted from a transmission circuit becomes M/f0. An ultrasound pulse of such pulse length is transmitted and received by way of a probe, during which time a received pulse dulls along its time-axis direction, that is, the depth direction, thus spreading, owing to frequency-dependent loss of an object and/or the band characteristic of the probe. Additionally, during the reception to display, the received signal undergoes processing using filters for a variety of types of processing. This filtering processing causes the waveform of the received pulse to dull, thus making it further spread in the time-axis direction.
By contrast, when the contrast agent is not used, the spreading is limited to a little amount. It is considered that the reason is as follows. A blood flow signal is low in intensity, which is an amount slightly larger than an apparatus noise level. The spread roots of a received pulse are mostly less than the noise level, while the remaining pulse portion of which strength is over the noise level keeps amounts slightly larger than the M/f0, or slightly lower that it, in some cases. Normally the display is carried out with gain adjusted so as not to visualize noise, with the result that the influences due to the spread pulse waveform scarcely appear on an displayed image. In the end, the blood vessel diameter in the depth direction in the image remains within a deterioration range of an original resolution, which is formed by adding the pulse length M/f0 to the original diameter itself of a blood vessel, and can be practically used.
However, when using the contrast agent, reflected signals from the contrast agent becomes a dominant in blood flow signals to be imaged, and their intensities increase by an amount of few dozens of decibels. As a result, most of the roots of a received pulse that has spread in the time-axis direction exceed the noise level, thus causing the blooming. This blooming generates in the azimuth direction as well.
As described above, there is a possibility that fine blood vessels can be seen through injecting the contrast agent. An important factor in observing blood flows passing through the fine blood vessels (significant blood flows are those of tumor and the coronaria) is to acquire realtime performance that assures the blood flow of being detected instantaneously at time when it began flowing, without fail. Since the conventional apparatus is obliged to observe blood flows of fine blood vessels in the CFM mode, a shortage of the number of frames becomes fatal. It is clear that a few frames in the CFM mode are still short for observation, and there is a fear that they fail to instantaneously detect a flowing blood.
Moreover, though injecting the contrast agent makes the detection sensitivity in the CFM mode improve up to a level as equivalent as the B mode, a conventional dynamic range for display in the power mode of the CFM mode is at most approx. 40 dB, which clearly shows there is a shortage in the range. Signals over 40 dB in sensitivity are all visualized with saturation at 40 dB, which results in poor gradation, thus providing only a blood flow image that gives a flat feeling in gradation. There is also a fear of degrading diagnosis capability on account of such poor gradation. Further, when diagnosing through observing the luminance levels of a blood flow image, the saturated portions cannot provide precise information about luminance levels.
In this way, where a blood flow is observed with the conventional apparatus with an objects into which the contrast agent is injected, there are various problems or drawbacks described above, thereby leading to a lowered performance in diagnosis, providing no practical use.
The foregoing conditions can be summarized as below, in which listed are representatives of unfavorable situations or unsolved problems about imaging blood flows using the conventional diagnostic ultrasound apparatus.
(1) In imaging blood flows of fine vessels on the conventional color Doppler technique, taking low-level blood flow signals into account, a variety of countermeasures to improve detection sensitivity are adopted, which include lengthening a transmission pulse length and various spatial average processing. But these countermeasures have been taken at the sacrifice of spatial resolution. Recently the basic performances of the apparatus have been advanced, and, like the improved power mode, there has been provided a mode to raise sensitivity without the sacrifice of the spatial resolution any more. The improved power mode is superior in sensitivity than the velocity mode of the CFM. Nevertheless, the power mode still succeeds a drawback that the resolution is low, like the conventional CFM mode. Primarily, without the heart""s cavities, blood flow images are desired to be used in observing the existence and/or running conditions of blood vessels, like seen in diagnosing tumor. Thus it is required to depict blood vessels including as finer vessels as possible.
Although such a demand has existed, a measure against the problem has not been proposed yet. Under such a demand, for example, U.S. Pat. Nos. 4,809,249 and 4,928,698 disclose a way of mapping an object in motion on the basis of a cross correlation method (correlation in the time domain) using blood flow velocities. However, an amount that can be detected using the cross correlation method is a blood flow velocity having poor sensitivity. Meanwhile, an apparatus capable of performing a blood flow mapping technique called color velocity imaging (CVI) has also been developed. This apparatus, however, not only uses an ultrasound pulse whose wave-continuous length is short, which is called short pulse, but also uses the cross correlation method, with color mapping of blood flow velocities inferior in sensitivity performed.
Because these techniques focus on obtaining velocities of objects in motion such as blood flow, there is a difficulty in sensitivity for detecting blood that flows at slower speeds or fine blood flows. Therefore, a recent demand for ultrasound diagnosis that the existence itself of such blood flow is desired to be observed and confirmed at a higher accuracy has not been met yet.
On one hand, in the case that blood flows of fine vessels are tried to be observed with the conventional diagnostic ultrasound apparatus with the contrast agent injected into an object, which has been eagerly developed recently, there are the following problems.
(2) Although enhancement effects of a contrast agent can greatly raise detection sensitivity, the spatial resolution of blood flow images is deadly deteriorated, making diagnosis difficult.
(3) When an object to be observed is flow of fine blood vessels, there is a possibility that an observer fails to see significant behaviors of blood flows on account of the shortage of realtime performance. In such a case, diagnostic capability decreases extremely, lowering reliability in diagnosis.
(4) When enhancement effects of a contrast agent increases a detection sensitivity of blood flow up to a level as high as the B mode, the dynamic range for display in the power mode belonging to the CFM mode might become short. In such a case, display gradation in a blood flow image is saturated, lowering reliability in display. Additionally there is a fear that information on luminance lacks, lessening diagnostic capability.
The present invention has been attempted to break through the foregoing situations the conventional techniques encounter. One object of the present invention is to provide a blood flow image which spatial resolution is raised, especially, which depicts fine blood vessels or blood streams flowing at slower speeds, with great fineness.
Another object of the present invention is to provide a blood flow image possessing improved spatial resolution and higher S/N, even when blood of fine vessels is observed.
Still another object of present invention is to provide, in imaging blood of fine vessels utilizing enhancement effects of echo signals from a contrast agent administered into an object, an blood flow image having improved image quality for the power mode as well as improved spatial resolution and superior realtime performance.
In order to realize the foregoing objects, a diagnostic ultrasound apparatus according to the present invention is based on, as one aspect, a configuration that comprises: scanning means for not only scanning a cross section of an object to be imaged by transmitting at least two times, along the same direction in the cross section, an ultrasound pulse having a wideband frequency characteristic, but also acquiring an electric echo signal caused by the ultrasound pulse every time of transmission; processing means for performing processing with a train of data of the echo signal aligning in a time axis direction at each sample position in the cross section and being acquired by the scanning means, the processing being for extracting a signal of a moving element; producing means for producing the signal processed by the processing means into two-dimensional image data; and visualizing means for visualizing an image based on the two-dimensional image data.
The processing means consist of means that perform desired processing for extracting a change in a moving element or in a phase with the train of data.
Preferably, the desired processing performed by the processing means highpass filtering or differential processing.
Still preferably, the two-dimensional image data are data indicative of luminance information or power information of the echo signal reflected by the moving element existing in the cross section, or data indicative of luminance information or power information of the echo signal originated from a contrast agent residing in the cross section.
On one hand, a method of ultrasound diagnosis according to the present invention is characterized in that comprising the steps of: for not only scanning a cross section of an object to be imaged by transmitting at least two times, along the same direction in the cross section, an ultrasound pulse having a wideband frequency characteristic, but also acquiring an electric echo signal caused by the ultrasound pulse every time of transmission; performing processing with a train of data the echo signal aligning in a time axis direction at each sample position in the cross section and being acquired through the scanning, the processing being highpass filtering or differential processing; producing the signal processed into two-dimensional image data; and visualizing an image based on the two-dimensional image data.
One example of operation of the present invention derived from the above-mentioned configurations will now be described.
An imaging mode on the present invention is to be referred to as a xe2x80x9chigh resolution flow mode.xe2x80x9d This xe2x80x9chigh resolution flow modexe2x80x9d allows high resolution color flow mapping images or grayscale flow mapping images, which represent existence or non-existence of blood flow, to be displayed by a contrast echo method or a non-contrast echo method.
Specifically, as an ultrasound pulse having a wideband frequency characteristic, i.e., an ultrasound pulse of which spatial resolution is high is transmitted and received along an object""s cross section in the same direction, a plurality of times, the cross section is scanned, producing a beam-formed echo signal. From a train of data along the time axis at each sample position in the scanned cross section, which are formed from the echo signal, unnecessary clutter compotes (reflected signal components stationary or almost stationary tissues) are removed, an echo component of blood flow (echo component from a contrast agent) being, extracted. This echo component is produced into data of appropriate forms (practically, data consisting of luminance or power), and displayed as a blood flow image.
Such a blood flow image is provided as an image indicative of luminance or power of a wideband echo signal. Hence, compared to a conventional CFM blood flow velocity image that represents blood flow velocities where a transmission pulse length is long, fine blood flows or flows of which speeds are slow are steadily detested as well, so that provided is a blood flow image which is higher in resolution and sensitivity and which finely represents existence of blood flow. Hence information about existence of blood flow can be provided with great reliability based on improved detectability of blood flow.
On one hand, an ultrasound pulse is set to a wideband pulse, with the result that the band of a received signal varies depth by depth because of the effect of signal""s frequency dependency within a living body. Accordingly, in response to the depth, the frequency of a reference signal for phase detection is controlled. Additionally, a filter whose band characteristic is variable, which is inserted into a receiving processor, is controlled in its band characteristic in response to the depth. Thanks to these configurations, both high resolution and high sensitivity are realized.
Moreover, injecting a contrast agent increases remarkably sensitivity. Thus, even if the number of transmission/reception times along the same direction in an object""s cross section is reduced than scanning with no contrast agent, there can still be provided image of enough sensitivity and quality. Especially, when a signal-processing filter for extracting a blood flow signal is formed by a differentiator, clutter removal is possible by a minimum of two times of transmission/reception. In addition, in the case that a technique of simultaneous reception in a plurality of directions toward transmission in one direction is adopted, enough sensitivity and quality of images are obtained, however the number of receiving directions may be increased. In consequence, using both the techniques of the reduced number of transmission/reception times and/or increasing the number of simultaneous receiving directions is able to provide blood flow images of very high frame rates (time resolution).
Further, because of highly increased sensitivity resulting from a contrast agent, a display dynamic range for the power of echo signals needs to be widened largely. Determining the display dynamic rage in agreement with this power can provide the high-quality blood flow images, without image saturation (i.e., lack of blood flow information).
A reflected echo from a contrast agent includes high-level harmonics as well as a fundamental wave. The present invention, regardless of frequencies of these reflected waves, can be applied to any of the fundamental wave, harmonics, and a mixed wave of those.
To remove unnecessary clutter components, a high-pass filtering technique, which is high in removal capability, or differential technique, of which realtime performance can be set highly, may be available.
Echo components from blood flow (echo components from a contrast agent) are possible to be formed into image data in an appropriate form. For example, they may be power image data produced using power mode processing or may be luminance image data produced using B-mode processing.
Due to the fact that a contrast agent collapses within a range of sound pressures used in an ordinary diagnosis or vibrates irregularly, the phase of reflected echo signals vary irregularly. Thus, however blood flow (contrast agent) may be nearly at rest in tissue, like blood perfusion, echo signals from the contrast agent, of which phases have varied, are extracted by highpass filtering or differential processing, differently from tissue echoes (clutters). Therefore, applying the present invention to an object into which a contrast agent makes it possible to detect through the fundamental wave the perfusion which was impossible to detect in the past, providing high-resolution, high-sensitivity, an realtime perfusion images.
The other constructions, operations and advantages according to the present invention will become distinct from the following embodiments and description on the accompanying drawings.